Detection of Biomolecules
The detection of organic molecules in extremely tiny amounts is challenging. Optical and electromagnetic properties describe the ability of the compound under investigation to absorb and alter a steady or alternating energy flux. Each compound does absorb light energy, absorbance. Some dissipate this energy as heat while others emit it partly in form of light of lower frequency, fluorescence. Light can be applied and collected by scanners with lateral resolutions down to micrometer range.
Molecular electromagnetic and electronic properties are in the limelight for sensor applications. To detect magnetic properties of electron or nucleus spins substantial amounts of material are needed. Electrochemical properties, reduction potentials, can be utilized to identify redox-active centers in hosts, but only a limited number of compounds carry such centers. Measurements of molecular conductivities are difficult through the challenges of establishing proper working connections.
Therefore, different methods have been developed to increase the conductivity of both, electrode-molecule connections and molecules themselves, e.g. through the metallization of polynucleic acids by adsorption and subsequent reduction of silver ions [Braun, 1998]. The utilization of conductivities as measurand in sensor array networks is cumbersome because each measurement cell needs to be contacted by two electrodes. Such setups require a potentially high number of electrodes and contacts respectively. In networks of interconnecting electrodes, where at each intersection or junction the conductivity can change due to analyte interaction, it is not possible to distinguish between the electrical current passage through one particular junction and alternative pathways thru other junctions. It is not possible to achieve matrix field or lateral resolution of m×n measurement cells with m×n electrodes when measuring resistances.
Without the need of a net direct current, DC, it is possible to probe a sample volume through a changing electric field, e.g. caused by an alternating current, AC. Electrical readouts provide in general the potential to use capacitive electric field, inductive magnetic field, resistive sensing and combinations thereof. Capacitive biosensors are disclosed in U.S. Pat. No. 5,532,128, US 2004/0110277 or WO 2009/003208. The general principle is the change of the dielectric hence capacitance in the sensor element through the presence of target molecules.
Capacitive or impedimetric biosensors are based on impedance spectroscopy, IS, electrochemical impedance spectroscopy, EIS, or charge based capacitance measurements, CBCM. An in-depth explanation of the mathematical derivation of the physico-chemical and electrochemical parameters concerning IS has been written by A. Lasia [Lasia, 1999]. An overview by Lisdat and Schäfer [Lisdat, 2008] focuses on the use of EIS in biosensing.
EIS is a complex method which measures simultaneously all elements of a circuitry including charge transfer processes at the interfaces and conductivity of the electrolyte. Impedimetric biosensors were tested for immunosensing, enzyme studies, cell based assays and nucleic acid determination. Focusing at the interfacial reactions, experiments with nucleic acids have shown that EIS is capable to discriminate between single stranded ssDNA, and double stranded dsDNA [Haso{hacek over (n)}, 2002; Strasák, 2002], to monitor hybridization [Drummond, 2003; Kelley, 1999; Fojta, 2003; Hashimoto, 1994; Pänke, 2007], melting of hybrids [Scuor, 2007] and intercalation [Xu, 2006]. Furthermore, EIS could detect single base pair mismatches [Cho, 2006; Pänke, 2008, Vermeeren, 2007], differentiate between DNA structures such as B- or M-DNA, probe DNA-analyte interactions, e.g. DNA with the cross-linking agent cis-platin [Yan, 2001] or with specific DNA binding proteins [Li, 2004].
Besides the monitoring of interfacial reactions it is also of great interest to detect materials like nucleic acids in the volume phase. The dielectric behaviour of DNA has been investigated over a wide frequency range from ultra low to microwave frequencies of 70 GHz. In contrast to water with a dielectric constant, ∈r, between 88 and 55.3 at 0 and 100° C., respectively, does a diluted solution of DNA display larger ∈r-values with approximately 120 at 20° and 100 kHz for 1% DNA solution in water [Takashima, 1984]. However, many pure organic polymers have smaller ∈r-value ranging typically between 2 and 10 at 20° C. Some polymers with widely extended electron orbitals can show ∈r-values as large as 105 as it has been measured for 2-chloroanthraquinone/tetra-chlorophthalic anhydride at 1 kHz [Pohl, 1985].
Because signal differences are usually small, increasing the signal to noise ratio is important. The classic approach uses large working electrodes which provide the significant capacitance in the system. Dendrimers [Li, 2007] and polymer structures are used to further maximize the surface areas [Maupas, 1997]. Peroxidase coupled reactions have been exploited as transducer [Ma, 2006] as much as nanoparticles [Peng, 2006; Li, 2005; Cai, 2003] or liposomes [Patolsky, 2003; Patolsky, 2001].
Alternatively, signal intensity can be enhanced through the alignment of the electrodes [Van Gerwen, 1998; Laureyn, 2000; Gheorghe, 2003; Dharuman, 2005] where the strength of the electric fields can be maximized by keeping electrodes in very close proximity [Montelius, 1995]. The signal increases linearly with length which can be realized through the winding structure of interdigitated electrodes, IDE. So far, distances depend on the manufacturing process and range from 10 to 1 μm [Brewood, 2008] with electrode widths as small as 500 to 200 nm [Van Gerwen, 1998]. Such IDE's have been used for investigating DNA hybridizations [Gheorghe, 2003; Dharuman, 2005; Van Gerwen, 1998; Hang, 2004; Berdat, 2006] and for the measurement of DNA concentrations. 200 pF responses have been obtained for 1 pM 2.961 bp phagemid pBluescript DNA solutions, which corresponds to approximately 1,1 DNA molecules per μm3 only [Henning, 2008]. The extrapolation to smaller electrode sizes hence smaller volumes and signals indicate the possibility to detect very few DNA molecules with signal differences in the order of 10 fF.
Further advancements can be envisaged through the use of nanogap-impedance sensors which promise high sensitivities due to their large surface-to-volume ratio and because electrode polarization effects become negligible. For minimizing shielding through buffer-ions, a microwave frequency, 1.28 GHz, has been used to sense antibody-thrombin binding in a 75 nm gap of an electrode area of 96 μm2 [Schlecht, 2006]. Nanocapacitors based on pore structures and gaps are discussed and modeled with the aim to use the method for polynucleotide sequencing [Sigalov, 2008; Lu, 2008]. These publications highlight the influence of the electric double layer, EDL, impedance which generates interfering noise levels at low frequencies kHz range. When the gap sizes become comparable or smaller as the EDL thickness, ddl, the nanogap capacitance will become independent of ionic strength, I. So is ddl 960 nm for I of 10−7 M but only 9.6 nm for 10−3 M solutions. In addition, the electric field distribution changes with gap size. Smaller gap sizes lead to potential increases in the centre and potential decreases at the electrode surfaces. The electric field becomes more homogeneous [Yi, 2005]. Eventually, for gaps, tubes, wires and other devices in the range of 10 nm and below quantum effects become important for the accurate description of reactions at electrode surfaces. So far, numerous techniques like electron beam lithography [Hwang, 2002], electrodeposition, electromigration [Iqbal, 2005] or other electrochemical methods, composite layer build-up combined with etching [Steinmüller-Nethyl, 2009] and fracture techniques [Reed, 1997; Reichert, 2002] have been employed to separate two electrodes by a tiny nanometer sized gap.
Segregation of Biomolecules and Sensor Arrays
In contrast to sequential measurements the quasi instantaneous measurement of molecules in complex mixtures requires segregation of the sample. This segregation reduces the degree of entropy in the system and facilitates efficient parallel measurements. Different classes of molecules are segregated into separated locations, e.g. spots. The predetermined position characterizes the class of molecules whereas the signal intensity at this position determines the amount which translates into concentrations of the compound in the sample. Microarrays are typical examples where single probes, sensor compounds or sequences, have been immobilized at a solid surface in specified regions [Southern, 1997]. For example, after a complex mixture of different sequences has been applied to a DNA microarray, only molecules with complementary sequences will hybridize to predetermined spots and generate a certain signal pattern. Disadvantages of such single probe designs are that many different classes of molecules might contain the same sequence, e.g. in the case of cognate genes or their splice variants, and each molecule might contain sequences which match to numerous different probes. The data analysis becomes very challenging and remains ambiguous.
Longer sequences also allow using of dual probes. Here, two single probes act like molecular brackets which are specific towards two distinct sites. The design allows more flexibility as one single bracket and can be chosen to target conserved regions, e.g. certain core exons of genes, while the other bracket is reaching to a region of flexibility, e.g. to exons which are characteristic for certain splice variants. Signal amplifications are possible, when the probes are designed as primers for polymerase chain reactions, PCR, which can be carried out in volume phase and at solid supported phase like sensor surfaces. Here, two primers are applied to a single sensor surface either sequentially or as a mixture to form a rather ideal 2D mixture of molecules at such surface. If a target molecule binds to such surface it can initiate an origin or seed for amplification and in succession a surface supported PCR reaction. Here, one difficulty is that different sites of the target molecule react with the very same surface. The target molecule and its copy, the amplicon, firstly, do not stick to the surface to enable the efficient enzyme catalyzed polymerization at the surface, but secondly, “bends” towards this surface in order to react with the second probe. Such amplification has been named “bridge amplification” because the molecules from a bridge from one probe to another probe [Boles, 2002]. The bridges develop along a single 2D surface where seeds can only grow geometrically to form small product islands which extend predominantly along their edges only. The resulting PCR efficiency, E, can start of high, with E being close to one. E will drop after several cycles [Mercier, 2003; Adessi, 2000].
The combination of impedimetric sensors and dual probes makes it visible to employ two differently functionalized electrodes in close proximity. The individual modification of electrodes is already difficult and becomes extremely challenging in nanometer range. The making and addressing or functionalizing of such electrode structures would require the use of e.g. very expensive electron-beam lithography techniques. In WO 2009/003208 Steinmüller-Nethyl et al. have proposed to use different materials for each electrode while the electrodes are separated trough an insulating layer with a thickness of only several nm. Here, different electrode materials enable the successive and selective binding of the molecular probes. However, the number of electrically conducting but different materials together with a specific and effective binding chemistry is limited and not practical for large arrays. The geometrical alignment in a set of crossing rows and columns is not possible. Importance of the publication is the description of the principle where analyte molecules with two selective binding sites bridge two electrodes containing one corresponding binding site respectively, and where subsequently the bridging molecules are detected not through DC but AC analytical methods. In US 2002/0022223 and US 2006/0019273 Connolly et al. have already described electrode couples where each electrode has to be modified with one type oligonucleotides. Analytes which contain corresponding sequences to both electrodes can hybridize, bridge and electrically connect electrodes. Such reactions can be recorded through DC signal changes. However, the selective modification of pre-manufactured electrode assemblies is very challenging, time consuming and expensive. Furthermore, all DC sensors are aligned in parallel. Because the number of electrodes and in particular their contacts increases linearly with the number of sensor pads such devices approach fast their technical limitations regarding integration density.
An identical approach has been followed in WO 2010/104479 using electrode arrays where the sensor action occurs at electrode crossings between electrode edges. The electrodes are just nanometers apart from each other and separated by an insulating layer. Here, the problem of selectively immobilizing capture probes on one of two corresponding electrodes across a separating step in the order of few nanometers, experimentally realized were 5-20 nm, has been recognized to be impossible by means of robotic spotters. The chosen method involves the binding of thiol-functionalized probes to all gold-electrodes, the selective removal by electrochemical stripping and repeated binding of thiol-functionalized probes to the second gold-electrode and so forth. Each functionalizing requires 2 hours for the binding step alone plus the time which is required for additional stripping and washing steps. Not only the production method is unsuitable to build complex sensor arrays, also the method of conductance measurements is preventing to utilize truly combinatorial approaches of the capture probes as described above. WO 2010/104479 presents a sensor array as the combination of different capture probes which are immobilized at parallel electrodes, e.g. rows, and bind to specific sequence in mRNA molecules, with explicitly one annealing probe at all columns. The chosen 21 nucleotide long capture probes are so long to serve the purpose to be specific for one particular mRNA only. This markedly high specificity has been chosen to detect per row only one target mRNA each. Therefore, the opposite annealing probe contains one single universal polyT-sequence which binds to all polyA-ends of the mRNA in the sample. Such array presents the complexity of just m×1. The additional columns only increase the effective sensor area by multiplying the number of identical sensors. The presented system is consistent with the chosen measurement method which only allows to distinguishing signals which arise from one entire line, here for example from one entire row.
The smallest dielectric gap concepts are not only aiming to quantify but also to sequence nucleic acids [Lee, 2005]. The proposed structures measure only a few nanometers. Those gaps can be described as tiny plate capacitors which record nucleotide specific changes of the dielectric as polynucleic acids pass through such gaps. The dielectrics properties of the nucleotides are one contribution, the other are the effects that the dielectric constant of bound and semi-bound water is significantly smaller than the one of free water. The estimates for the reading speed are based on using MHz frequencies and would therefore range in the order of 1 Mio reads per second.
The integration of individual sensors is essential for detecting different compounds simultaneously in one sample. Early attempts dealt with difficult sensor arrays where each field had been connected though a separate pair of electrodes [Albers, 1999]. In WO 2004/001405 Frey et al. describe a design and operation of a biosensor array where multiple biosensor fields are arranged matrix-like on a substrate. Each field is addressed through one actuator and one detector line, and each line is able to address several fields. At the time of probing one particular field all other lines are set to a fixed potential e.g. floating ground, unless fields were grouped before. Although it is not explicitly stated, the description contains interdigitated electrodes IDE in each sensor field. The design solved the problem of addressing many fields with minimal connections, however the degree of integration is limited and the manufacturing technologies of said structures are expensive. The issue of modifying said structures differently in each sensor field remains challenging when it comes to tiny dimensions. Along the same line are proposals made by Maeda [2004], Maracas G. [2000] and Li [Li C., 2005], who are separating individual sensor test cells. A similar approach has been followed in CN 101046458 by Liu [2007] who describes an array of crossing electrodes on a substrate which have been insulated from their neighbours through separated micro flow ponds.
Further, from the state of the art the publications WO 2010/1204479 and EP 2088430 are known. WO 2010/1204479 is directed to a sensor for detecting a nucleic acid molecule comprising an electrode arrangement with two electrodes and nucleic acid probes immobilized at the surface of the electrodes. The present invention also refers to a kit and a method of using the sensor or a sensor array. The present invention is further directed to a process of manufacturing a sensor and sensor array.
EP 2088430 provides a bio-sensor including nanochannel-integrated 3-dimensional metallic nanowire gap electrodes, a manufacturing method thereof, and a bio-disk system comprising the bio-sensor. The bio-sensor includes an upper substrate block having a plurality of metallic nanowires formed on a lower surface thereof and including an injection port through which a biomaterial-containing sample is injected, a lower substrate block having a plurality of metallic nanowires formed on an upper surface thereof, and a supporting unit supporting the upper and lower substrate blocks so that the upper and lower substrate blocks can be disposed spaced apart at a predetermined distance to form a nanochannel, wherein the metallic nanowires formed on the upper and lower substrate blocks are combined to form three-dimensional metallic nanowire gap electrodes.